Magnetic resonance (MR) imaging is a known technology that can produce images of the inside of an examination subject without radiation exposure. In a typical MR imaging procedure, the subject is positioned in a strong, static, homogeneous base magnetic field B0 (having a field strength that is typically between about 0.5 Tesla and 3 Tesla) in an MR apparatus, so that the subject's nuclear spins become oriented along the base magnetic field. Radio-frequency (RF) excitation pulses are directed into the examination subject to excite nuclear magnetic resonances, and subsequent relaxation of the excited nuclear magnetic resonances can generate RF signals. Rapidly switched magnetic gradient fields can be superimposed on the base magnetic field, in various orientations, to provide spatial coding of the RF signal data. The RF signal data can be detected and used to reconstruct images of the examination subject. For example, the acquired RF signal data are typically digitized and stored as complex numerical values in a k-space matrix. An associated MR image can be reconstructed from the k-space matrix populated with such values using a multi-dimensional Fourier transformation.
Magnetic resonance can be used to produce images representing tissue perfusion, which refers to the delivery of oxygen and nutrients to tissues of a subject by means of blood flow. Perfusion studies allow an assessment to be made of in vivo organ functions such as, e.g., brain activity. In certain MR perfusion imaging techniques, a contrast agent that generates a signal detectable by magnetic resonance imaging is injected into a subject. Magnetic resonance image data are acquired at a time when the contrast agent has optimally flowed into the region or anatomy of interest. Since the contrast agent is injected into the vascular system of the subject, the appearance of the contrast agent in the magnetic resonance image is representative of blood flow in the region or anatomy of interest. Magnetic resonance perfusion techniques are particularly useful in the context of magnetic resonance images of the head, in particular the brain, wherein cerebral blood flow (CBF) is identified.
Another MR perfusion imaging technique uses “arterial spin labeling” (ASL) instead of injection of a contrast agent. In an ASL procedure, a spatially selective inversion or saturation of water protons in arterial blood is used to label or “tag” blood flowing into the region to be imaged. For brain perfusion imaging, this tagged region can be a slab in a lower portion of the head and/or neck area below the region to be imaged, from where blood flows up into the brain. When the labeled or tagged blood reaches the tissue within the imaging region, it attenuates the MR signal emanating from the perfused tissue following spatially-selective excitation of the region. “Subtraction” of a labeled image from a control image (i.e., one obtained without labeled/tagged blood in the imaged region) can provide a measure of the amount of tagged blood that flowed into the imaged tissue. This quantity is closely related to the local tissue perfusion. The image subtraction can be achieved, e.g., by performing a voxel-by-voxel subtraction of image intensity between a tagged image and an immediately preceding or subsequent control image. In ASL procedures, the imaged region or volume will thus have different magnetization histories arising from the tagged and control pulse sequences.
The difference in the MR signal intensity for labeled and control images is typically only a few percent of the tissue MR signal, and thus MR images based on ASL differences often suffer from the influence of image noise. In some ASL imaging procedures, many repetitions (e.g. 10-50) of the ASL data acquisitions are averaged to increase the signal-to-noise ratio (SNR). Such factors of imaging time and small signals can make ASL techniques prone to motion corruption and artifacts, which can accumulate when acquiring data for a plurality of images. ASL techniques and certain approaches for reducing noise effects are described, e.g., in U.S. Pat. No. 8,203,340 of Pfeuffer.
The use of 3D encoding in ASL imaging, e.g., using a 3D gradient and spin echo technique (GRASE), can provide higher SNR than conventional slice-by-slice methods. The higher SNR can be based on, e.g., the faster echo refocusing larger number of echoes in an echo train compared to conventional EPI techniques. Further, such 3D sequences facilitate simultaneous readout that can better image a particular extent of tagged blood perfusion over a region of interest as compared to a 2D approach, where slices are encoded sequentially and thus at different perfusion times. The use of 3D encoding in ASL imaging procedures is described, e.g., in D. A. Feinberg et al., Cerebral Blood Flow Imaging With 3D GRASE ASL Sequence Increases SNR and Shortens Acquisition Time, MAGNETOM Flash, 62-69 (March 2009). However, such high-resolution 3D MR imaging can still be susceptible to motion corruption. Thus, the potential benefits of a 3D ASL imaging procedure, for example, in evaluating stroke in a subject, can be mitigated by its sensitivity to patient head motion due to the associated segmented k-space acquisition and subtraction artifacts arising from successive acquisition of control and labeled (tagged) images within the scan.
So-called “navigator” sequences are additional RF pulses that can be used in MR imaging procedures to dynamically track anatomical motion. Navigator pulses are typically spin echo (SE) or gradient echo (GRE) sequences. Echo signals returned by a navigator can be used to correct for motion during certain types of image acquisition sequences. For example, the use of volumetric (3D) navigator sequences to correct for motion effects is described in M. D. Tisdall et al., Volumetric Navigators For Prospective Motion Correction And Selective Reacquisition In Neuroanatomical MRI, Magnetic Resonance in Medicine, 68:389-99 (2012). The approach of Tisdall et al. is limited to imaging pulse sequences where the magnetization history is the same for every acquired navigator, and when motion is detected the entire imaging volume is adapted (i.e. translated and/or rotated) accordingly, to compensate for the detected change in position. Such a navigator approach is not compatible with typical 3D GRASE sequences, in which the magnetization history changes depending on whether a particular sequence is in the labeling or control phase. For example, in the labeling phase of a 3D GRASE ASL sequence, an additional tagging step is inserted in the sequence. When motion occurs during the imaging procedure, the imaged volume should be adapted to compensate for the motion, whereas the tagged volume should not be adapted because it is typically below the brain in the neck where the major arteries can be labeled.
Accordingly, it would be desirable to have a system and method for imaging perfusion that addresses some of the shortcomings described above, including effective motion correction that can be done in real time.